Prosthetic heart valves are used as a replacement for natural heart valves of patients. A standard implantable mechanical heart valve typically includes an annular valve housing or body (often called an "orifice") to provide a lumen or passageway therethrough for blood flow. One or more occluders mounted to the valve are movable between an open position, allowing blood flow, and a closed position which blocks blood flow. In many mechanical valves, the occluders are essentially plate-like members called "leaflets." Typical configurations include one, two or three leaflets in the valve body.
An attachment mechanism typically surrounds the valve body and is used to secure, typically with sutures, the valve to the patient's heart tissue. While some early prosthetic valves used hooks or barbs for attachment, a fabric suture or sewing cuff which is secured to the annular valve housing is typically used. Attachment of the suture cuff to the valve may be through any of a number of different retention techniques, some of which provide rotatable coupling. For example, U.S. Pat. No. 5,360,014 shows a separate stiffening ring which carries a suture cuff and which is clipped to the valve body by a lock wire between the valve body and the stiffening ring.
There has been an ongoing effort to improve the efficiency of prosthetic heart valves. One critical factor in heart valve efficiency is the total area of the lumen when the leaflets are in an open position. For patients with small aortic roots (typically defined as a tissue annulus diameter of between about 17 mm and about 21 mm), there have been indications that available prosthetic valves are stenotic when compared to the healthy native valve. The orifice or lumen area of typical prosthetic valves is so small that the left ventricle may be unduly burdened in maintaining an adequate cardiac output. The effective orifice area is further reduced by the hydrodynamic impedance of the valve. It has been found that currently available small prosthetic aortic valves are associated with decreased tolerance to exercise, reduced rate of regression of left ventricular hypertrophy and a higher incidence rate of congestive heart failure. (See "Prosthetic Valves for the Small Aortic Root," Journal of Cardiac Surgery, 1994; 9[suppl]: 154-157, by H. B. Barner, A. J. Labovitz and A. C. Fiore.)
One technique which provides a less stenotic replacement valve involves enlargement of the aortic root and tissue annulus by the surgeon. However, such procedures introduce additional risk to the patient because they require greater manipulation and excision of tissue. Further, these procedures require an increased duration of heart-lung bypass, thereby imposing additional risks to the patient from that procedure. Another surgical approach for implanting a less stenotic valve has been to implant tissue valves such as allografts and stentless heterografts in these patients. However, for many patients, the well-established durability of mechanical heart valves is preferred.
To meet the need for less stenotic small prosthetic heart valves, changes in mechanical valve sewing cuff configurations have been introduced. This has allowed implantation of valves having a lumen diameter typically one size (2 mm) larger than has been previously possible. For example, the tissue annulus of the standard mechanical heart valve from St. Jude Medical, Inc., of St. Paul, Minn., lies on sewing cuff fabric which extends from a pyrolytic carbon orifice ring. In the Hemodynamic Plus (HP) Series mechanical heart valve also available from St. Jude Medical, Inc., the sewing cuff lies entirely between cuff retaining rims of the orifice ring so that the cuff is implanted supra-annularly and the upstream retaining rim periphery or circumference constitutes the valve surface (the "valve tissue annulus") engaging or apposing the heart's tissue annulus which remains after excision of the native valve. The intra-annular and subannular projection of this valve reduces the potential for tissue overgrowth of the valving mechanism and maintains the patency of the valve and tissue lumens.
Another prior art prosthetic heart valve is depicted in U.S. Pat. No. 5,360,041, issued Nov. 1, 1994. In this configuration, the valve is completely supra-annular. The suture cuff forms a brim which surrounds the extreme edge of the upstream annulus of the orifice ring. Although this may allow for increased valve and lumen size, the high supra-annular profile of the valve has, in at least some patients, blocked the right coronary ostium. Further, the position of the suture cuff may render the valving mechanism relatively vulnerable to tissue overgrowth. In addition, there is no intra-annular barrier to retard growth of tissue into the valve lumen.
While recent developments in prosthetic heart valves, such as those described above, have provided improvements, they remain stenotic compared to the healthy native valve. Improvements to further decrease the transvalvular pressure gradients of forward blood flow would be beneficial to patients. Although small, non-stenotic replacement valves are typically needed for the aortic position, there is also a need for such valves for the mitral position, typically in pediatric cases.
Another problem which may be associated with replacement heart valves with small lumens relates to formation of thrombus and thromboembolism. Thrombus and thromboembolism are known complications of mechanical heart valves and can result in serious disability or death. To help prevent these complications, a common treatment involves life-long anticoagulant therapy. However, anticoagulant therapy itself leads to an increased risk of anticoagulant-related hemorrhage.
Factors which influence the risk of thrombus and thromboembolism formation for mechanical heart valve patients include the nonphysiological surfaces and blood flow introduced by mechanical valves. Further, typical mechanical heart valves subject the blood to high shear stress, largely because the relatively small lumens of such valves tend to produce high velocity forward flow as the heart strives to maintain adequate cardiac output. Since the blood flow velocity immediately adjacent to the walls of the valve lumen and the occluders must be zero, large velocity gradients are generated during forward flow as a consequence of the high mean velocity. The shear stresses are proportional to the velocity gradients. High shear stresses are known to activate blood platelets and damage red blood cells. Such damaged red blood cells release a biochemical agent, adenosine 5'-diphosphate (ADP), which further activates platelets. The activated platelets have the potential to be deposited on the valve or downstream from the valve and to aggregate into thrombi. Furthermore, the activated platelets and the released biochemical agents initiate a coagulation cascade. Therefore, valves with mean forward flow velocities and peak shear stresses which are lower than prior art valves would be beneficial to patients.